This application claims the benefit of a priority under 35 U.S.C. 119 to Great Britain Patent Application No. GB 0423322.7 filed Oct. 20, 2004, the entire contents of which are hereby incorporated by reference.
This invention relates to a coil for a magnet and to a method of manufacturing a coil for a magnet. More particularly, it relates to a gradient coil and to a method of manufacturing a gradient coil, and in particular, to a gradient coil for use in a magnetic resonance imaging (MRI) system.
MRI systems are used today for investigating a large variety of body parts. These systems are based on nuclear phenomena displayed by atomic nuclei having a non-zero magnetic moment (or “spin”). When such nuclei are placed in a static, uniform magnetic field, the nuclear spins are aligned by the magnetic field so as to be either aligned with or against the static magnetic field. The nuclear spins are not stationary, but precess around an axis defined by the magnetic field. The frequency at which the spins precess is known as the “Larmor frequency” ω0. The Larmor frequency is given by:ω0=γB0 where γ is the gyromagnetic ratio of the nucleus and B0 is the applied magnetic field. For a hydrogen nucleus, for example, γ=42.57 MHz/T.
When the nuclear spins are aligned in the static magnetic field B0, it is possible to “flip” the spins by applying an alternating magnetic field B1. In order to do this, the alternating magnetic field must be at 90° to the static magnetic field and it must alternate at substantially the Larmor frequency. When such an alternating field B1 is applied, the spins will tend to align themselves parallel to B1, and when the alternating field is removed, the spins will relax back into the state in which they are aligned with the static magnetic field B0. The alignment of the spins with the alternating field decreases the magnetization in the longitudinal direction (parallel to B0) and increases the magnetization in the transverse plane (that is, the plane perpendicular to B0), and the subsequent relaxation of the spins when the alternating field is removed produces the reverse effects. These changes in the magnetization are detected in the MRI process, and are processed to provide a visible display of the nuclei.
FIG. 1 at 11 shows a typical MRI system in block diagram form. The magnet 12 provides the static magnetic field B0. In principle, the magnet 12 could be a superconductive magnet, an electro-magnet or a permanent magnet. However, a super-conducting magnet is commonly used, since these readily provide a large, homogeneous static magnetic field. The magnet 12 contains a bore 13 enabling the entry of a patient into the static magnetic field. A patient shown at 14 is inserted into the bore 13 using a bed arrangement 16 so as to be within the static magnetic field.
Radio frequency (rf) pulses generated by transmitter 22 and applied through multiplexer 23 and radio frequency coil apparatus 24 act to tip the aligned spins so as to have a projection, for example, in the X-Z plane; the X-Y plane or the Y-Z plane. The X, Y, and Z nomenclature refers to the imaginary orthogonal axes shown at 21 used in describing MRI systems; where the Z axis is an axis co-axial with the axis of the bore hole. The Y axis is the vertical axis extending from the center of the magnetic field and the X axis is the corresponding horizontal axis orthogonal to the other axes.
The spins when realigning after the radio frequency pulse is removed generate free induction decay (FID) signals which are received by the radio frequency coil apparatus 24 and transmitted through the multiplexer 23 to the receiving circuit 26. From the receiving circuit the received signals go through the controller 25 to an image processor 27. The image processor works in conjunction with a display memory 28 to provide the image displayed on display monitor 29. It should be noted that the radio frequency coil apparatus 24 can comprise separate coils for transmitting and receiving or the same coil apparatus 24 could be used for both transmitting and receiving the rf pulses.
In order to spatially resolve the MRI signal, encoding signals within the static magnetic field are provided by gradient coils (not shown in FIG. 1). There are typically three sets of gradient coils. X gradient coils alter the strength of the Z component of the static magnetic field along the X axis, Y gradient coils alter the strength of the Z component of the static magnetic field along the Y axis, and Z gradient coils alter the strength of the Z component of the static magnetic field along the Z axis. The strength of the Z component of the static magnetic field in other directions, than the X and Z axes for example, can be changed using two or three of the gradient coils in combination.
The X, Y and Z gradient coils are driven by X gradient driver 17, Y gradient driver 18 and Z gradient driver 19, respectively. It is possible to modify the local static magnetic field B0, at a particular point in space using the gradient coils so that only nuclei within a small volume element of the patient have a Larmor frequency equal to the frequency of the rf field B1. This means that the F.D.I. signal comes only from nuclei within that volume element of the patient. In practice the gradient coils are supplied with time-varying electrical currents from a power supply, such as a power amplifier, so that the volume element in which the nuclei have a Larmor frequency equal to the frequency of the applied rf field scanned over the patient so as to build up a 2-D or 3-D image of the patient.
A typical prior art set of gradient coils is disclosed in, for example, “Foundations of Medical Imaging” by Z. H. Cho et al. (published by Wiley International), and is shown schematically in FIG. 2. The X gradient coils are shown in FIG. 2(a). FIGS. 2(b) and 2(c) show the Y gradient coils and the Z gradient coils, respectively.
It is common practice to provide an actively screened gradient coil, which comprises an inner cylindrical assembly, and an outer assembly disposed coaxially and concentrically with respect to the inner. The outer assembly is connected in series opposition to the inner assembly and the composite design is chosen to reduce the external field produced by the whole, which would create unwanted eddy-current effects in the structure of the magnet. Typically there will be fewer turns on the outer assembly than on the inner assembly.
It will be noted that the X gradient coils and the Y gradient coils shown in FIGS. 2(a) and 2(b) are in the form of saddle coils. In each case, two saddle coils are placed on either side of the X-Y plane.
In the prior art, the gradient coils are constructed over a tubular base. In one possible arrangement, the X gradient coils are disposed over the tubular base, the Z gradient coils are placed over the X gradient coils, and finally the Y gradient coils are placed over the Z gradient coils (although the order in which the gradient coils are provided on the former is not limited to this particular order).
An X gradient field may be generated by utilizing a set of at least four X gradient coils A, B, C, D with appropriate current senses laid on a first cylinder. (See FIG. 3 of the accompanying drawings). The four gradient coils may be termed saddle coils, as discussed above. They have one or more planes of symmetry (three in the case of an X or Y gradient coil). The four gradient coils (saddles) may be connected in series. A Y gradient coil resembles an X gradient coil, rotated through 90° about the Z axis. X and Y gradient coils may be manufactured by cutting or etching a complex track in a sheet of electrically conductive material. To minimize heat dissipation, it is desirable to leave as much conductive material in the sheet as possible, resulting in a pattern having conductive tracks of variable widths.
The X and Y gradients currently available are made from a flat copper plate 30 (typically 2 mm in thickness as depicted in FIG. 4) with a single ‘spiral’ cut or track 32 in the copper plate 30 to form the electrical circuit. Four such plates are arranged as quadrants (e.g., A, B, C, D) on the surface of a cylindrical form 34 to create the whole ‘X’ gradient, for example, as shown in FIG. 3. The resultant tracks 32 are wide and relatively small in number in order to match the available power supplies. This results in several undesirable affects. The small number of turns results in quantization effects in which the design can vary only by single full turns. However, the relative difference between N and N+1 turns can be quite large where N is small. The linearity, screening or strength, therefore, cannot all be matched simultaneously to be within a certain accuracy. An additional disadvantage is that the connection between each of the quadrants requires an additional connecting conductor which is disposed on either a top or a bottom of the copper plate, which takes up valuable build space. A third disadvantage is that the wide tracks result in an uncertain current path during rapid pulsing. In particular, eddy-current effects cause the current to run at the extremities of the tracks that in effect manifests as a non-constant resistance versus frequency of an applied ac current. The resulting image has time-dependent distortions that must be corrected.
In some gradient assemblies of prior art, attempts have been made to drive different parts of the circuit with separate power supplies. For example, coils A and D of FIG. 2a are driven by one power supply and coils B and C are driven by a second power supply, or alternatively, coils A and B are driven by one power supply while coils C and D are driven by another. Thus, each power supply need only provide a smaller output power to achieve the same gradient strength, and/or the gradient can be drive more rapidly for the same peak voltage. Therefore, the power supplies will be smaller, less expensive and more reliable. One serious disadvantage of this arrangement is that in order for the gradient field to be accurately linear, as required by in imaging equipment, it is necessary that the current from each is accurately identical at all times during the pulse sequence. This results in unachievable requirements for maintaining accurately identical currents with identical phase at all times from the power-supply/gradient combination and such attempts have largely been abandoned.